Implant

ABSTRACT

The disclosure lies in the field of regenerative medicine and relates to an implant having a matrix material, and a method for manufacturing an implant having matrix material. The disclosure further relates to a crimped implant.

BACKGROUND

The present disclosure relates to an implant comprising a matrixmaterial, and a method for manufacturing an implant comprising matrixmaterial. The disclosure further relates to a crimped implant.

A relatively new field of medicine—since the early 1990s—is the field ofRegenerative Medicine. Regenerative Medicine is the process of creatingliving and functional tissues to repair, replace, or restore tissue ororgan structure and function lost due to age, disease, damage, orcongenital defects. This field of medicine uses new methods including(stem) cell therapy, development of medical devices and tissueengineering.

Over recent years, continuous improvements in our healthcare haveresulted in dramatic demographic changes, e.g. an increase in theaverage age of the population. These demographic changes are causing anincrease in the prevalence of diseases associated with aging, such ascardiovascular diseases. Many of these diseases arise from the loss ordysfunction of specific cell types in the human body, leading topermanently damaged tissues and organs.

Cardiovascular diseases are one of the biggest causes of deathsworldwide. One way to treat at least some of these diseases is by tissueengineering. Tissue engineering can be used for the replacement ofcardiovascular tissues, such as arteries and heart valves. Currentlyused cardiovascular substitutes encounter risks due to coagulation,infections, degeneration, and no growth possibilities. Tissueengineering uses patient's own cells and a biodegradable polymerscaffold to make autologous tissue that is able to grow, adapt andrepair. To ensure proper cell and tissue growth, the scaffolds must behighly porous and match the mechanical properties of the tissue.Electrospinning is a technique that produces polymer nanofibers using ahigh voltage electrostatic field. It results in a highly porous materialconsisting of nanofibers that resembles the extra cellular matrix of thetissue. Tissue engineering can, for example, be used for coronary bypassgrafts, heart valve replacements, AV shunts for dialysis patients.

In the field of surgery, minimal invasive surgery is preferred. Howevertissue engineered constructs can often only be implanted via normalsurgical procedures since the constructs cannot be compressed to a sizesufficiently small to facilitate minimal invasive surgery. Someartificial heart valves can now be crimped to a diameter of 18 French (6mm) to allow for implantation via a small peripheral incision (e.g.transfemoral or transjugular). However, many elderly patients in need ofa replacement valves also suffer from stenotic and therefore narrowedarteries, which currently excludes them from the much preferred minimalinvasive surgery. A reduction of 1 or 2 French in crimpable diameteralready means a significant increase in the number of treatablepatients.

The technique of tissue engineering consists of constructing substitutes(e.g. biological substitutes) for diseased tissues. Tissue engineeringmakes use of natural or polymeric scaffolds that provide mechanicalsupport and promote the re-growth of cells lost due to trauma ordisease. A scaffold is a temporary structure used to support material(e.g. tissue) during the recovery thereof.

Polymeric scaffolds can be constructed from biocompatible, non-toxicpolymers. The choice of polymer and the technique used to make thescaffold effects the mechanical properties exhibited by the scaffold.

In the publication of Bouten et al, Advanced Drug Delivery Reviews,2011, vol. 63, pp 221-241 synthetic polymers have been demonstrated tobe good substrates for valvular and vascular tissue engineering. Forcardiac tissue engineering, the most commonly used biodegradablesynthetic scaffold materials are polyglycolic acid (PGA), polylacticacid (PLA), polyhydroxybutyrates (PHB), ε-polycaprolactone (PCL) ortheir co-polymers. No functioning implants were disclosed using thedescribed synthetic scaffold materials.

In the publication by Dankers et al in Nature Materials, 2005, Vol. 4,pp 568-574 solution cast polymer films comprising2-ureido-4[1H]-pyrmidinone (Upy) polymers were shown to be non-toxicwhen studied in vivo. However the use of UPy polymers as cardiovascularimplant scaffolds was not shown.

In order to obtain a tissue engineered construct, a scaffold can bepre-seeded in vitro with the appropriate cells prior to implantation. Inmost cases, as the formation and the remodeling of the newly formedtissue proceeds, degradation of the scaffold should slowly and steadilytake place, leaving only new healthy tissue behind By “degradation,” itis meant the breakdown of the material into smaller parts, e.g. chemicalcompounds and/or elements that can be eliminated from the body by meansof excretion in urine for example.

A drawback of growing a tissue construct in vitro is that the completeprocedure including growing and implanting has to be conducted sterilelymaking it a costly and laborious procedure. In addition, regulatoryguidelines on living tissues are complex, resulting in long and costlyprocesses towards product approval.

Another option is to seed an implant with cells prior to implantation.This method requires the harvesting of cells from the subject to receivethe implant, optionally growing the cells in vitro, and seeding thecells in the construct followed by implantation. This method has thesame down sides as the previously described method.

BRIEF SUMMARY

It is an object of the present disclosure to provide for an implantwhich can be implanted via minimal invasive surgery.

It is an object of the present disclosure to provide for an implantwhich regenerates tissue in vivo.

It is further an object of the present disclosure to overcome one ormore of the above mentioned draw backs associated with the prior art.

One or more of the above objects have been reached by the embodiments ofthe present disclosure. The inventors have surprisingly found that theabove objects are reached with an implant comprising one or moresupramolecular compounds and wherein the matrix material comprises afibrous network and having a porosity of at least 60%, preferably aporosity of between 70% and 90%.

The embodiments are illustrated in more detail in the followingdescription and with reference to the drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows native fiber orientations of a heart valve leaflet (Sauren(1981)).

FIG. 2 shows helical fiber orientation in blood vessel (Holzapfel, J.Elast., 2000).

FIG. 3 shows results of valve testing at 120/80 mmHg on a PCL-bisureavalve at systemic conditions showing stable performance for 20 hours.

FIG. 4 shows pictures of PCL-bisurea valves at start (two upperpictures) and after 20 hours (two lower pictures) of systemic conditionsin a valve test, 20 hours (R11020).

FIG. 5 shows results of valve testing at 50/25 mmHg on a PCL valve atsystemic conditions showing a decreasing in performance within severalhours.

FIG. 6 shows pictures of PCL valves at start (two upper pictures) andafter 20 hours (two lower pictures) at 50/25 mmHg in a valve test, 20hours (R11020).

FIG. 7 shows Photograph (a) and schematic overview (b) of set-up forvalve testing.

FIG. 8 A shows the uniaxial tensile test results for PCL and PCLbisureaelectrospun scaffolds along and perpendicular to (FIG. 8 B) the mainfiber direction.

FIG. 9 A shows the results of the uniaxial fatigue tests (10%, 2 Hz,) onPCL and PCLbisurea. FIG. 9 B shows comparative results of conduitfatigue testing conducted using PCLbisurea and PCL UPy implants.

FIG. 10 shows the results of the crimping tests on electrospun heartvalve implants.

FIG. 11 shows the different crimped states of 3D electrospun heartvalves.

FIG. 12 shows the stents after different stages of crimping for 30electrospun heart valves.

FIG. 13 shows three pictures of DAPI staining of a PCL-bisurea implantin an ovine model.

FIG. 14 shows two SEM images of PCL and PCL-bisurea matrices.

DETAILED DESCRIPTION

In the present description and appended claims, the following terms areused, which are explained below.

A “polymer” is intended to also include homopolymer, copolymer orsupramolecular polymer unless otherwise stated.

A “supramolecular polymer” is a polymeric array of monomeric units thatare brought together by reversible and highly directional secondaryinteractions, resulting in polymeric properties in dilute andconcentrated solutions, as well as in the bulk. The monomeric units ofthe supramolecular polymers themselves do not possess a repetition ofchemical fragments. The directionality and strength of thesupramolecular bonding are important features of these systems, that canbe regarded as polymers and behave according to well-establishedtheories of polymer physics.

A “supramolecular monomer compound” in this application is a compoundwhich by virtue of reversible and highly directional secondaryinteractions (with other supramolecular monomer compounds) forms asupramolecular polymer.

A supramolecular polymer is therefore composed of monomers, whichmonomers are designed in such a way that they autonomously self-assembleinto the desired polymeric structure. This is in contrast toconventional polymerization reactions whereby monomers are linked viacovalent bonds. As a result of the self-assembly, a (much) highervirtual molecular mass of the material is achieved. Examples ofsupramolecular polymers have been described in, for example, Science,1997, 278, 1601.

A “contrast agent” is a substance used to enhance the contrast ofstructures or fluids within the body in medical imaging.

A “scaffold” is a temporary structure used to support material (e.g.tissue) during the formation and/or recovery of said material.

A “structural component” is the part of the scaffold that is intendedfor the provision of structural properties.

An “imaging component” is the part of the scaffold that is intended forthe provision of imaging properties.

A “biologically active component” is the part of the scaffold that isintended for the provisional of biological activity.

A “substrate” is the material on which growing of cells takes place.

An “implant” is a medical device that can replace a dysfunctional ordamaged biological structure, support a damaged biological structure,cover a damaged biological structure, or enhance an existing biologicalstructure.

“Backbone of a polymer” is the backbone chain (also called main chain)and is a series of covalently bonded atoms that together create thecontinuous chain of the polymer.

“Porosity” is measured, for example, by mercury porosimetry, fluidintrusion and gravimetry.

“Pore size” is the average size of openings (pores) in the matrixmaterial. Porosity as mentioned in the specification is measured asfollows:

The scaffold weight is measured using a balance. The dimensions (lengthand thickness for tubes, length, width and thickness for sheets) arealso measured. The porosity is calculated using the following formula:

porosity=(1−Density of scaffold/density of polymer)×100%

-   -   where density of polymer varies based on the polymer used and        density of scaffold is calculated as weight of scaffold/volume        of scaffold.

By “pores,” it is meant the inter-fiber spaces (that is, the pore size)in the matrix material. The pore size and porosity are properties of ascaffold that influence the attachment, proliferation, migration and/ordifferentiation of cells.

A “minimally invasive procedure” is a procedure (surgical or otherwise)that is less invasive than open surgery.

The inventors have surprisingly found that the above object is reachedwith an implant comprising a matrix material having a porosity of atleast 60%. Said implant can by implanted in a subject (also calledrecipient of the implant) with a minimally invasive procedure. Due tothe porosity of at least 60%, the implant can be compressed minimizingthe size of the implant (also referred to as crimping). Since the sizeof the implant is reduced, a smaller opening is required forimplantation resulting in less discomfort to the recipient of theimplant and minimizing recovery time of the recipient. Preferably, theporosity is between 70% and 90%. Further a fiber structure with such aporosity allows for diffusion of nutrients into the matrix and for theingrowth and/or infiltration of cells into the matrix.

The inventors have found that they can solve one of the majorlimitations currently associated with minimal invasive implantation ofe.g. prosthetic heart valves by allowing crimping to a smaller sizecompared to existing transcatheter heart valves.

In an embodiment of the present disclosure, the implant is acardiovascular implant and preferably an implant chosen from the groupconsisting of a (blood) vessel, a heart valve, a cardiovascular patch ora valved conduit. It is beneficial for the recipient that an implantaccording to the present embodiment can be implanted by minimal invasivesurgery. For these implants, only a minor incision has to be made inorder to facilitate implantation. Preferably, the implant is applied toan object (i.e. a patient) via a small incision. Due to the highporosity of the implant the implant can be reduced at least 5 times indiameter size going from fully expanded towards fully crimpedconfiguration and back.

In an embodiment of the present disclosure, the implant is reinforced byat least one support structure, preferably the at least one supportstructure is chosen from the group consisting of a reinforcement ring, asuture ring or a stent structure, and is preferably biodegradable,preferably the implant consists of the reinforced matrix material. Thepresence of the support structure can, for example, be aimed atreinforcing the implant, allowing crimping of the implant during aminimally invasive procedure, allowing fixation of the implant at thecorrect anatomical position or allowing repeated puncturing of theimplant with a needle. Suitable support structures are those widelyknown in the art and are, for example, used in artificial heart valvesor as coronary stents or stents for other kinds of arteries.Reinforcements are for example described in U.S. Pat. No. 4,626,255,U.S. Pat. No. 6,338,740, US 200410148018, U.S. Pat. No. 3,570,014 andU.S. Pat. No. 4,084,268.

In an embodiment of the present disclosure, the implant has a matrixmaterial which consists of a fibrous network. Said fibrous network ismade up out of fibers. The fibers allow for the implant to have goodstructural integrity while maintaining its porosity and pores.Preferably, the fibrous network are electro spun fibers.Electro-spinning is a technique using a metal target or mold, havingeither a flat, plate-like form or a complex three-dimensional form,depending of the preform that is desired. Polymer fibers are depositedonto this mold by means of an electromagnetic field. The polymer fibersare generated from a solution of one or more polymers in one or moresolvents. This technique of electro-spinning is known in the art andwill not be further in detail in this specification. In Dutch patent NL1026076 (corresponding to US 2008/0131965), the preparation of anarticle by means of electro-spinning of polymer microfibers isdisclosed. The electrospinning setup used in the development of such aproduct is climate controlled and also allows for control of spinningarea, nozzle speed, rotation of the collector and the possibility of theapplying a small negative voltage (up to −4 kV). The humidity,temperature and other above mentioned factors can be used alone or incombination to alter various characteristics of the electrospun fibers.These include, but may not be limited to, fiber morphology, fiberdiameter, fiber and pore size distribution, porosity and scaffoldthickness.

In an embodiment of the present disclosure the fibers of the matrixmaterial consists of one or more supramolecular polymers. By using thesekinds of polymers, the inventors have found that the implant can beimplanted in a subject without having to seed the implant with cellsprior to implantation. By using supramolecular compounds, the inventorsobserved that cells attach, infiltrate and grow in vivo on and in theimplant while the implant fulfills the function of the tissue to bereplaced or repaired. Hence the implant can be implanted directly in thepatient. The advantages of implanting a scaffold directly are areduction in production time and cost, the scaffold can be stored forprolonged periods of time and it qualifies as a medical device meaningthat regulatory approval can be obtained much faster. Hence, one or moreof the important drawbacks of growing a tissue construct, or seedingcells, in vitro are overcome, while keeping the advantage and promise ofregeneration, leaving only new healthy tissue behind.

In an embodiment of the present disclosure, the implant isbiodegradable. This allows for the implant to be degraded after beingimplanted in the body. Hence the implant is replaced over time bytissue. The advantage is that the implant does not have to be removedsurgically preventing further discomfort for the patient who receivedthe implant.

The inventors found that if one or more supramolecular compounds have abackbone comprising or consisting of polycaprolactone (PCL) or acombination of PCL, caprolactone, poly lacticacid and/or lactic acidvery good results are obtained with respect to the structuralcharacteristics of the implant and with respect to biodegradability.

In an embodiment of the present disclosure, the implant has a matrixmaterial that comprises one or more supramolecular compounds wherein theone or more supramolecular compounds further comprise one or more groupschosen from Upy (ureido-pyrimidinone) and/or bisurea. A polycaprolactonepolymer containing UPy-groups, is for example disclosed in patentapplication EP 1 687 378. Preferably, the supramolecular compound isPCL-bisurea (also referred to as PCLbu). The present inventors havefound that with these compounds good results can be obtained withrespect to the in vivo growth of tissue in the implant as well as forstructural characteristics of the implant. Good results are obtainedwith PCL-bisurea.

In an embodiment, the one or more supramolecular compounds comprises atleast PCL UPy, preferably the matrix material consists of PCL UPy.Especially good results are obtained with PCL-UPy.

A method for preparing PCL bis urea can be found in, for example,“Biomaterials by the supramolecular control of nanofibers by E. WisseISBN: 978-90-386-1094-8 chapter 2. The chemical formula of PCL-bis ureawith a butyl spacer is shown in the Formula I below. Alternatively, nospacer or another alkyl spacer, e.g a hexyl spacer can also be used. ThePCL acts as a soft-block while urea groups make up the hard block. Thehydrogen bonding between the hard blocks results in reversible physicallinks.

Wherein p and n are integers and p and n>1. The p and n values can bealtered to yield different formulations of PCL bis urea which mightresult in different properties. The p value is dependent on the startingmolecular weight of polycaprolactonediol while the n value relates tothe number of chain extensions of the PCL bis urea. Preferably values ofn can range from 4-40.

A method for preparing PCL Upy can be found in, for example,“Biomaterials by the supramolecular control of nanofibers by E. WisseISBN: 978-90-386-1094-8 chapter 6. PCL-UPy polymers can be prepared, forexample, as comprising the urea hydrogen bonding group (Upy-U 1) or theurethane hydrogen bonding group (Upy-U 2).

In an embodiment of the present disclosure, the fibrous networkcomprises nano fibers and/or micro fibers, preferably the diameter ofthe micro fibers ranges from 3 to 20 micrometer and preferably from 5 to10 micrometer. The advantage of this diameter is the excellentmechanical and structural stability of the implant, while providing amicrostructure that is sufficiently porous to allow for cell ingrowth(Balguid, Strategies to optimize engineered tissue towards native humanaortic valves, PhD thesis Eindhoven University of Technology, 2008, ISBN978-90-386-1185-3).

Nano fibers are fibers having a diameter of less 1 micrometer. Thediameter of micro and nano fibers can be obtained by measuring thediameter under a microscope.

In an embodiment of the present disclosure, the matrix materialcomprises pores having a diameter ranging from 1-300 micrometer andpreferably ranging from 5-100 micrometer. The advantage of these poresizes is that it allows the passage of the cells to be cultured andhence a good infiltration of cells into the complete thickness of thepreform, which is required to ensure formation of tissue throughout thecomplete preform. The requirements of the pore size depend on the sizeof the cells to be cultured and can be selected according to this size.The size of human cells is generally larger than the size of animalcells, hence the differentiation between the most preferred pore sizeswhen using either animal or human cells.

In a preferred embodiment of the present disclosure, the matrix materialforms a layer having a thickness of at least 100 μm and maximally 3000μm, preferably the thickness is between 200 and 1000 μm. In case theimplant consists of matrix material, the implant has a thickness of atleast 100 μm and maximally 3000 μm, preferably the thickness is between200 and 1000 μm. The inventors have found that with the definedthicknesses implants are obtained with good structural properties sothat these implants can fulfill the desired function when implanted in asubject while resulting in tissue growth and obtaining tissue of goodquality.

In an embodiment of the present disclosure, the implant has a linearelastic stiffness ranging from 0.1-50 MPa and preferably from 0.1-10MPa. The inventors have found that these ranges provide the combinationof strength and flexibility that is required to withstand thehemodynamics of the human cardiovascular system. Previously publishedresults on native and tissue engineered materials report stiffnessvalues in this range as well (Non-invasive assessment of leafletdeformation and mechanical properties in heart valve tissue engineeringby Kortsmit, ISBN: 978-90-386-2002-2 (2009). Stradins et al. (2004),Clark, (1973)).

In an embodiment of the present disclosure, the implant has an (linear)elastic regime of at least 30%, and preferably at least 45% and morepreferred at least 60% (Linear elastic stiffness has been measured usingstandard uniaxial tensile tests (for a description see ISO 13934-1:1999Textiles—Tensile properties of fabrics—Part 1: Determination of maximumforce and elongation at maximum force using the strip method)). Thisensures that an implant according to the present embodiment does notshow plastic deformation or breakage within a physiological strainregion (strain exerted in the body on the implant). For example, fornative heart valves, physiological strains have been reported of ˜60%(Billiar & Sacks, (2000), Driessen et al., 2005)). Current bioprostheticvalves already fail to comply with these values, showing maximal strainsduring diastolic pressurization of gluteraldehyde treated porcine valvesof 2-4% and 3-10% in circumferential and radial direction, respectively(Adamczyk & Vesely, 2002), and an average strain value of 4-10% in theheart valve leaflets (Sun et al., 2005). Furthermore, chemically fixedanisotropic tissue was described to become more isotropic (Zioupos etal., 1994) and less compliant than fresh tissue (Broom et al., 1982;Schoen et al., 1997; Billiar & Sacks, 2000) due to chemicalcrosslinking. Commercially available polymers that were shown to have ashort elastic regime in a uniaxial tensile test, also showedinsufficient performance in an in vitro valve testing set-up. Thus, forthe implant to be used in the body and fulfill the function of thetissue to be replaced or repaired an extended elastic regime isimportant.

In an embodiment of the present disclosure, the implant, the fibers (ofthe matrix material) have a preferred orientation direction. Preferably,the fibers in the implant are arranged in such a way that when theimplant is implanted, the fibers are arranged substantiallyperpendicular to the blood stream.

Preferably, the preferred fiber alignment is circumferential around animaginary axis of the implant wherein the axis points in the directionof bloodflow in case of a tubular implant. In case of a heart valveleaflet, the fibers are preferably arranged in the same fashion asdepicted in FIG. 1.

Such an orientation can be introduced during manufacturing of theimplant (i.e. with electro spinning). Such a fiber structure mimics thenatural fiber alignment in native tissues, e.g. the hammock-likecollagen architecture in native heart valve (as depicted in FIG. 1.Sauren et al. 1981) and the helical collagen fiber orientation in nativearteries (as depicted in FIG. 2, by Holzapfel (2000))

By mimicking the extra cellular matrix of the natural environment, atissue can be grown having good structural properties, which eventuallydevelop towards a native-like architecture.

In an embodiment of the present disclosure, the linear elastic stiffnessratio between the stiffness in the preferred fiber direction and thestiffness perpendicular to the preferred fiber direction is at least2:1, preferably at least 4:1 and more preferred at least 10:1 and evenmore preferred at least 50:1. Implants having such a ratio have goodstructural properties while still providing for a substrate for cells togrow on mimicking the natural environment.

In an embodiment of the present disclosure, the implant furthercomprises biologically active compounds and/or contrast agent. In orderto monitor how fast the in vivo degradation of the implant proceeds, andin order to judge the ultimate success of the tissue engineeringprocedure, a contrast agent can be present in the implant. Said contrastagent can be visible in relevant clinical imaging techniques, such ascomputed tomography (CT), magnetic resonance imaging (MRI) and/ordiagnostic sonography (or ultrasonography) using ultrasound for imagingpurposes, for example. A preferred contrast agent is described inEuropean patent application having application Ser. No. 10/193,654, afluorinated polymer having a glass transition temperature (Tg) below 40°C., preferably below 20° C., more preferably below 0° C., as an imaginglabel or contrast agent in 19F magnetic resonance imaging (MRI). Theamount of fluorine (19F) in the fluorinated polymer is preferably atleast 5 wt %, based on the total mass of the polymer. Said fluorinatedpolymer comprises at least one polymer selected from the groupconsisting of (per)fluorinated polyethers, (per)fluorinated polyesters,(per)fluorinated poly(meth)acrylates, and (per)fluorinatedpolysilicones, preferably (per)fluoroethers. Said polymers can beincorporated in the polymers making up the fibrous network.Alternatively the polymers are present separately in the fibrousnetwork.

Biological active compounds can be added to, for example, promote cellinfiltration, retainment, differentiation and proliferation, as well astissue formation and remodeling.

The present disclosure further relates to a method for manufacturing animplant, preferably as described above, having a matrix material one ormore supramolecular compounds, wherein the matrix material has 60%porosity and preferably between 70 and 90% porosity, comprising thesteps of:

-   -   Providing a mold;    -   Applying the matrix material to the mold by means of electro        spinning of one or more supramolecular compounds; and    -   Separating the matrix material from said mold.

The obtained implant has the same advantages as described above. In anembodiment of the present disclosure, the method further comprises thestep of providing at least one support structure to the implant ormatrix material.

The disclosure further relates to a crimped implant, preferably theimplant according to embodiments as previously described, which can becrimped to a diameter size of up to 20% compared to the diameter of theimplant before crimping (uncrimped). This allows for the implant to beeasily provided by minimal invasive surgery. A further advantage is thatsince the implant can be crimped to 20% of its initial diameter size, itcan be provided to subjects that currently are excluded from receivingan implant via minimal invasive surgery because, for example, theirarteries are too narrow for a crimped implant according to the prior artto pass through. Preferably, the crimped implant is a heart valveimplant. Crimping an implant according to the present disclosure isachieved by methods known in the art. An example of crimping is given inthe examples.

The present disclosure further relates to a method for growing a valve,comprising the step of providing an implant, preferably an implantaccording to the present disclosure to a subject (a patient). Theimplant is preferably an implant according to the present disclosure.This step can be preceded by the step of making an incision in the skinof the subject.

After implanting the implant in the patient (human or animal) theimplant is capable of functioning (as the tissue to be grown) afterimplantation and before cellular ingrowth occurs, and wherein cellularingrowth occurs after implantation and the matrix material is degradedover time.

The embodiments will be further illustrated by the following nonlimiting examples. The appended claims also form part of the descriptionof the present application.

EXAMPLES Example 1

Implants were manufactured according to the following method. A requiredamount of PCL, PCL bisurea or PCL UPy is dissolved in an appropriatesolvent/solvent mixture and stirred until dissolution. The resultingsolution is delivered at a constant flow rate (flows varying with timeare also possible and will result in scaffolds with differentproperties) to a nozzle that can be electrically charged. Typically,this is done using a syringe pump. A high voltage is applied to thenozzle. Voltage differences (combination of positive and negativevoltages) used range from 10-20 kV, although it is possible to producefibers at other voltages. A rotating collector, usually in the form of acylinder is placed. The collector is connected to the ground or anegative terminal. The speed of rotation is typically 100 rpm. Thefibers are deposited on the collector. The length and thickness of theproduced implant is affected by flow, voltage, collector rotation speed,sweep and the nozzle speed. After the desired thickness is achieved, thespinning is stopped and the collector is removed. The implant in thecollector is vacuum dried and annealed (˜37° C.) overnight. The implantis removed by the collector by soaking it in warm water (˜37° C.),although other methods of removal are possible. A SEM image of anelectrospun fiber mesh is shown in FIG. 14.

Example 2

The implants obtained according to the method of example 1 were testedin accordance with ISO5840:2005. The hemodynamic performance of avalvular implant is assessed by loading with physiologically relevantflow and pressures by means of a mock loop system. FIG. 7 shows an imageof the mock loop used. The valves to be mounted in this mock loop aresubjected either to pulmonary pressures and flow or to aortic pressuresand flow. The fluid used in the experiments is a physiological salinesolution. In the mock loop system shown above, the circulation fluid isdisplaced by means of a computer (PC) controlled piston pump (P). Thepiston is connected to a servomotor (SM) system that is controlled by amotion control board. The piston fills the left ventricular cavity (LV;VLV=1 L) from a reservoir through a model mitral valve (MV) andsubsequently ejects the fluid through the arterial valve (AV) into aWindKessel (WK) model that consists of two resistances (R1 and R2) and acompliance tank (C; VC=2 L). From the Windkessel model, the fluid flowsback into the reservoir through a section of silicone tubing. Thepulsatile flow through the arterial valve is measured by a Flow Meter(FM1). Additionally, the mean Cardiac Output is registered by means of aClamp-On flow sensor (FM2) on the outlet silicon tube of the WindKessel.Ventricular and arterial pressure is recorded using pressure transducers(Pv and Pa) that are connected to a bridge amplifier (Picas. PeekelInstruments). The signals are recorded by a Data-Acquisition board andstored on a hard disk in a PC. An endoscope is introduced into thesystemic artery to which a high-speed color camera is connected (M5) tocapture valve dynamics throughout the heart cycle at a frame rate of 200Hz. Valvular pressure and flows are monitored over time to assess valvefunctionality.

Valves made from PCLbu and PCL were subjected to systemic conditions at120/80 mmHg for 20 hours. The obtained results clearly show betterresults for PCLbu (FIGS. 3 and 4.) implants than for PCL implants (FIGS.5 and 6).

The pictures of FIGS. 4 and 6 show an valve in the opened (left toppicture) and closed (right top picture) configuration at the start ofthe test and in the opened (left bottom picture) and closed (rightbottom picture) after 20 hours. The pictures of the tested valvesclearly show that the PCL valves are damaged whereas the PCLbu valvesremained largely undamaged.

Example 3 Uniaxial Tensile Test

The Zwick tensile stage is connected to a measurement module, which isconnected to a PC. After the necessary preparations, this system can beused to retrieve the stress-strain characteristics of strips of forexample biological tissue. First, both sides of the sample are mountedin the tensile stage with the use of the two clamps. Software on the PCtriggers the tensile stage, thereby elongating the sample. During thiselongation, displacement, as well as force, is recorded, andsubsequently stored in an TRA file. From this file, complemented withthe dimensions of the specimen, mechanical parameters (Young's Modulus,ultimate tensile strength/strain) can be determined using Matlab. RT(dry), grip-to-grip separation=9 mm, elongation rate=9 mm/min, 20Nloadcell.

Width and thickness are used to calculate the Young's Modulus [Pa],which is a measure for the stiffness of the strips.

$\begin{matrix}{E = \frac{{F_{2}l_{2}} - {F_{1}l_{1}}}{\left( {l_{2} - l_{1}} \right)A_{0\;}}} & \left( {{Eq}.\mspace{14mu} 1} \right)\end{matrix}$

F [N] is the force and l [m] the length of the sample (length[mm]*10⁻³).

A0 [(m²] denotes the cross-sectional area of the sample before testing(width [mm]*thickness [mm]*10⁻⁸).

FIG. 8 A shows uniaxial tensile test results for PCL and PCL-bisureaelectrospun scaffolds along (left) and perpendicular to (right) (FIG. 8B) the main fiber direction, showing a clear difference in elasticregime (straight line) between the 2 materials. For PCL plasticdeformation around 10% strain and breakage observed below 15% strain(perpendicular). For PCL-bisurea, no plastic deformation or breakagewithin physiological strain region (<60%) was observed under the sameconditions.

Example 4 Fatigue Testing A) Uniaxial Fatigue Test

Using a set-up similar to the uniaxial tensile test, cyclic loading wasperformed to assess the fatigue behavior in a uniaxial fatigue tests(10% strain, 2 Hz,) on PCL and PCLbisurea. Results confirm failure modesand timelines observed in the valve test (Example 2), showing fatiguewithin 1000 cycles (<% hour in-vivo) for valves made from PCL, and nofatigue for valves made from PCL-bisurea. Even after 1 million cycles(>11 days in-vivo) no fatigue damage was observed for the PCLbisureavalves (FIG. 9 A).

B) Conduit Fatigue Test

The implants obtained according to the method of example 1 were testedfor fatigue resistance by applying cyclic pressure loads to tubularshaped scaffolds. This benchtop test is used to determine the durabilityof a vascular device by subjecting it to hydrodynamic pulsatile loadingwith a relevant pressure differential. Device specimens are submerged inan environmental chamber and fatigued for 3 million cycles. Testconditions are designed to meet the requirements of in vitro mechanicalfatigue testing stated in ASTM F 2477-07 “Standard Test Methods for invitro Pulsatile Durability Testing of Vascular Devices” and FDA Guidance1545 (2010) “Non-Clinical Engineering Tests and Recommended Labeling forIntravascular Devices and Associated Delivery Systems”. The test iscontrolled and monitored via the pressure control method, which dictatesthat the test will control the cycle pressure range within desiredranges. The frequency was set to 5 Hz, temperature to 37° C., and thepressure ranges were set at 35/15 mmHg. As output, the outer diameter ofthe tested scaffold is measured as a function of the amount of cycles.

FIG. 9B clearly shows that PCLbu based materials have a higher increasein outer diameter as a function of amount of cycles when compared to UPymaterials, indicating that UPy materials are even more resistant tofatigue failure. Comparing the results from shown in FIG. 9 A with FIG.9 B it is clear to see that the PCL-bisurea withstands fatigue betterPCL implants and furthermore PCL-UPy implants withstand fatigue evenbetter then PCL-bisurea implants.

FIG. 15 also shows results for 2 different PCLbu configurations and 2different UPy configurations, which demonstrates that by tuning thescaffold material, the fatigue properties can be improved for PCLbu aswell as UPy based scaffolds

Example 5

Youngs modulus for isotropic and anisotropic electrospun PCLbu stripswere obtained from which the stiffness ratios were calculated. Theresults are listed in Table 1.

TABLE 1 Initial Youngs 1-6% strain modulus [MPa] n = 3 ParallelPerpendicular Ratio Isotropic  5.1 ± 0.29  4.2 ± 0.23 1.2:1 Anisotropic 31 ± 4.0 0.60 ± 0.10  52:1

Table 1 clearly shows that anisotropic spun PCLbu strips show a highYoung modulus when measured parallel to the preferred orientation of thefibers and a low Young modulus when measured perpendicular to thepreferred fiber orientation. Table 1 shows that for isotropic spun PCLbustrips the Young modulus measured in the parallel direction iscomparable to the Young modulus in the perpendicular direction.

Example 6 Crimping Test

Implants according to the present embodiment having a diameter of 28 mmwere subjected to a crimping test. The implants were shrunk to a certaindiameter using a commercially available device for crimpingtranscatheter heart valves. The different stages of the crimping processare depicted in FIG. 11. After that, they were allowed to return totheir normal size. After unfolding, the diameters were measured again.The results are shown in FIG. 10 wherein the striped bars represent thediameter to which the implant is crimped, the white bar represents thediameter of the implant after it has been allowed to return to itsuncrimped size and the black bar represents the same as the white barbut after soaking the implant in PBS at 37° C. for 1 hour.

From the results, it is clear that even after the implant has beencrimped to a diameter of 8 mm, return to its initial size (diameter of29 mm) was still observed after soaking in PBS at 37° C. for 1 hour.FIG. 12 shows pictures of the valves after being allowed to return totheir uncrimped size. The numbers next to the picture correspond to thenumbers of FIG. 10.

Example 7 In Vivo Testing

An implant according to the present embodiment was implanted at thepulmonary valve position in an adult sheep using open heart surgery.Cells were visualized with DAPI staining and cell infiltration wasassessed under a microscope. Cell infiltration was assessed 1 day, 7days and 8 weeks after implantation. The results are depicted in FIG.13. The upper left picture shows cell infiltration after 1 day and theupper right picture after 7 days for the PCLbisurea implant and thelower picture shows cell infiltration in the PCL-Upy implant after 8weeks. The pictures clearly show that cells in vivo infiltrate thePCLbisurea implant and the PCL-UPy implant.

1. An implant comprising a matrix material comprising one or moresupramolecular compounds, wherein the matrix material comprises afibrous network and the matrix material has at least 60% porosity. 2.The implant according to claim 1, wherein the implant is acardiovascular implant selected from the group consisting of a bloodvessel, a heart valve, a cardiovascular patch and a valved conduit. 3.The implant according to claim 1, wherein the implant is reinforced byat least one support structure selected from the group consisting of areinforcement ring, a suture ring and a stent structure.
 4. The implantaccording to claim 3, wherein the matrix material is biodegradable. 5.The implant according to claim 1, wherein the matrix material consistsof a fibrous network.
 6. The implant according to claim 1, wherein thematrix material consists of one or more supramolecular compounds.
 7. Theimplant according to claim 1, wherein the one or more supramolecularcompounds have a backbone comprising polycaprolactone (PCL), or acombination of PCL, caprolactone, poly lacticacid and/or lactic acid. 8.The implant according to claim 7, wherein the supramolecular compoundsfurther comprise one or more groups chosen from ureido-pyrimidinone(Upy) and/or bisurea.
 9. The implant according to claim 8, wherein theone or more surpramolecular compounds comprises at least PCL UPy. 10.The implant according to claim 5, wherein the fibrous network comprisesnano fibers and/or micro fibers.
 11. The implant according to claim 1,wherein the matrix material comprises pores having a diameter rangingfrom 1-300 micrometers.
 12. The implant according claim 1, wherein thematrix material has a linear elastic stiffness ranging from 0.1-50 MPa.13. The implant according to claim 1, a linear elastic regime is atleast 30%.
 14. The implant according to claim 5, wherein the fibers havea preferred orientation direction.
 15. The implant according to claim14, wherein the linear elastic stiffness ratio between the preferredfiber direction and perpendicular to the preferred fiber direction is atleast 2:1.
 16. The implant according to claim 1, wherein the implantfurther comprises biologically active compounds and/or a contrast agent.17. A method for manufacturing an implant comprising a matrix materialcomprising one or more supramolecular compounds, wherein the matrixmaterial has at least 60% porosity, comprising: applying the matrixmaterial to a mold by electro spinning of the one or more supramolecularcompounds to obtain an implant; and separating the implant and the mold18. The method according to claim 17, wherein the method furthercomprise providing at least one support structure in the implant.
 19. Acrimped implant that is obtained by crimping an implant according toclaim
 1. 20. The crimped implant according to claim 19, wherein thecrimped implant has a diameter size which is 20% or less compared to thevolume of the uncrimped implant.
 21. A method for growing a tissue in apatient, comprising implanting the implant according to claim 1 in thepatient.